This invention is generally related to apparatus and methods for measuring parameters associated with the flow of blood through a body segment. More particularly, this invention is related to apparatus and methods for monitoring cardiac output. Still more particularly, this invention is related to noninvasive apparatus and methods for continuously monitoring cardiac parameters.
Cardiac output is the volume of blood which the heart pumps in one minute and is one of the most important cardiovascular parameters. The cardiac output reflects the supply of oxygen and nutrients to tissue. Measurements of cardiac output provide invaluable clinical information for quanitifying the extent of cardiac dysfunction, indicating the optimal course of therapy, managing patient progress, and establishing check points for rehabilitation in a patient with a damaged or diseased heart, or one in whom fluid status control is essential. Exercise as well as pathological conditions of the heart and circulatory system will alter cardiac output; therefore, the measurement of cardiac output is useful both in rehabilitation and in critically ill patients.
Instrumentation currently in use for invasive and noninvasive measurement of cardiac output has several disadvantages. Cardiac output may be measured either invasively or noninvasively. The invasive techniques for measuring cardiac output involve penetration of the skin by a catheter, require complex instrumentation which must be operated by skilled personnel, and present a risk to the patient. Invasive techniques such as indicator dilution and thermal dilution allow only intermittent measurement of cardiac output since it is possible to obtain only one determination of cardiac output per injection in dilution methods.
The noninvasive techniques for measuring cardiac parameters include ballistocardiography, electrical impedance measurements, ultrasonics, phonocardiography and vibrocardiography. The instrumentation involved in present noninvasive techniques for measuring cardiac output is complex, expensive, inconvenient to use and requires highly trained operators. Existing electrical impedance instrumentation permits determination of cardiac output only during voluntary apnea, which means that the patient must hold his breath. Therefore, existing electrical impedance instrumentation is unsuitable for use with critically ill or unconscious patients.
Electrical bioimpedance measurements permit quanitifaction of blood flow as a result of changes in electrical conductivity of a body segment. The electrical impedance technique for measuring cardiac output is based upon changes in thoracic electrical impedance caused by cardiovascular activity. The change in impedance has several origins:
1. Cardiovascular activity causes pulsatile impedance changes. Approximately one-half of the pulsatile impedance change is related to volumetric changes of blood in the arteries as a result of arterial pressure compliance because blood is the most electrically conductive substance in the body. The other half of the pulsatile impedance change is caused by variation of the specific resistivity of blood as a function of blood velocity, which is related to the alignment of red blood cells.
2. Ventilation causes pulsatile impedance changes. In the thoracic region, the ventilation related impedance changes are directly proportional to the varying amount of air in the lungs while in the extremities the ventilation related impedance changes are caused by variation of the venus pool as a result of ventilation.
3. Edema and blood pooling cause nonpulsatile impedance changes.
The impedance changes listed as (1) above enter into the calculation for quanitification of cardiac output. The ventilation impedance changes (2) represent an unwanted background signal which must be suppressed to obtain accurate results.
The electrical impedance technique for measuring cardiac output has a significant advantage over the dilution methods, in addition to being noninvasive, in that the result of the measurement of the impedance change is the stroke volume, which is the volume of blood pumped per one heart contraction, which, when multiplied by heart rate, results in cardiac output. Therefore, stroke volume provides the clinician with beat-two-beat information on cardiac function. Indicator dilution methods determine average cardiac output only. However, the electrical impedance measurement technique has failed to gain wide acceptance since it has demonstrated a good correlation with invasive techniques only on healthy individuals. A strong deviation from expected values of cardiac output occurred in patients with pulmonary apnea or increased thoracic fluid content. The deviation is related to the mathematical structure of the equations used to calculate the stroke volume.
The first useful equation to quantify stroke volume from the cardiovascular impedance change was described by Nyboer, electrical impedance plethysmography, Charles C. Thomas, 1959. Myboer's equation is: ##EQU1## which uses the systolic downstroke extrapolation method on a graph of the thoracic impedance as a function of time to determine the cardiovascular impedance change during a heartbeat. R is the resistivity of blood of the subject, L is the average distance between sensing electrodes attached to the subject, .DELTA.Z is the cardiovascular impedance change, and Z.sub.o is the base electrical impedance of the thorax. Typical magnitudes of the thoracic variables for a healthy subject are Z.sub.o =30 ohms, R=150 ohm centimeters, L=30 centimeters, and .DELTA.Z=30 milliohms. Since ventilation impedance changes are typically an order of magnitude larger than the cardiovascular impedance changes, all measurements using Nyboer's method have to be performed during voluntary apnea. Therefore, Nyboer's method is currently used in peripheral applications only.
Kubicek et al, "The Minnesota Impedance Cardiography-Theory and Applications", Biochem. Eng. 9: 410, 1974, describe an improved equation for the calculation of stroke volume. Kubicek introduced the product of two cardiovascular variables, ventricular ejection time T and the maximum rate of impedance change .DELTA.Z/sec, in place of a single cardiovascular variable, .DELTA.Z used in Nyboer's equation. The maximum rate of impedance change can be derived either from the graphical extrapolation of systolic upstroke impedance change to maintain a rate of impedance change in one second or from the maximum value of electronically derived first derivative of impedance change. Kubicek's equation is: ##EQU2## where SV=stroke volume, R is the specific resistivity of blood, L is the distance between sensing electrodes, Z.sub.o is the base electrical impedance, T is the ventricular ejection time, and .DELTA.Z/sec is the maximum rate of impedance change.
In spite of the clear advantage of being capable of noninvasive cardiac output determination, electrical bioimpedance measurements have not gained wide spread use in thoracic measurements for the following reasons:
1. Good correlation against other techniques could be demonstrated only on healthy individuals;
2. Calculated cardiac output of patients with increased thoracic fluid content was much higher than the expected value;
3. There are unresolved questions about the influence of hematocrit on the accuracy of calculation of cardiac output;
4. The determination of cardiac output could be performed only intermittently with the subject in voluntary apnea;
5. The electrodes used with prior art devices are not suitable for long term monitoring; and
6. The existing instrumentation is high in cost and very complex, requiring skilled operators.
Therefore, there is a need for a system for monitoring cardiac output parameters which may be used continuously, is simple to operate, is safe for both the subject and operators, has the versatility to measure and display several cardiovascular variables for every heartbeat and which has relatively low initial cost and low operating costs.